System and Method for Automatic Tube Potential Selection for Dose Reduction in Medical Imaging

ABSTRACT

A method for CT imaging that utilizes an automatic tube potential selection for individual subjects and diagnostic tasks. The method quantifies the relative radiation dose of different tube potentials for achieving a specific image quality. This allows the selection of a tube potential that provides a reduced radiation dose while still providing CT images of a sufficient quality.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is based on, claims the benefit of, andincorporates herein by reference, U.S. Provisional Patent ApplicationSer. No. 61/225,818, entitled “METHOD FOR AUTOMATIC kV SELECTION FORDOSE REDUCTION IN CT,” and filed Jul. 15, 2009.

BACKGROUND OF THE INVENTION

The field of the invention is x-ray computed tomography (CT) and, inparticular, the selection of CT system parameters for reducing radiationdose.

X-ray imaging exposes individuals to ionizing radiation while beingimaged. This radiation dose has become an important concern for publichealth. While CT has played a dramatic role in the detection and stagingof disease, subjects are at a small but increased risk ofradiation-induced malignancy. The most frequently employed method toreduce radiation dose is to lower the x-ray tube current using subjectsize or weight-based protocols. Another important and related dosereduction technique is automatic exposure control (AEC), which involvesthe automatic adaptation of the tube current during the CT scan. Thatis, as the CT gantry rotates around the subject and as the subjecttravels through the gantry, the tube current is adjusted according tosubject size to keep image noise and quality constant. While thiswidely-used approach can achieve a dose reduction of 40-50% withoutsacrificing image quality, there are still other considerations thatcould lead to further dose optimization.

Recently, there have been several physics and clinical studies on theuse of lower tube potential, which is generally measured usingpeak-kilovoltage, denoted “kV” or “kVp”, CT imaging, with the purpose ofimproving image quality or further reducing radiation dose. A principlebehind this is that iodine has increased attenuation, or CT contrast, atlower tube potentials than at higher tube potentials in the range oftube potentials available on CT scanners. In many CT exams usingiodinated contrast media, to achieve the superior enhancement of iodineat lower tube potentials, improves the conspicuity of hypervascular orhypovascular pathologies owing to the differential distribution ofiodine, for example, in renal and hepatic masses, and inflamed bowelsegments.

Images obtained at lower tube potentials tend to be noisier, primarilydue to the higher absorption of low-energy photons by the subject.Therefore, a tradeoff exists between image noise and contrastenhancement in determining the clinical value of lower tube potential.When subject size is above a particular threshold, the benefit of theimproved contrast enhancement is negated by the increased noise level.In this situation, the lower tube potential may not generate betterimage quality than the higher tube potential for the same radiationdose. In other words, dose reduction may not be achieved with the lowertube potential because the higher tube potential is needed to maintainappropriate image quality. However, below this size threshold, variousdegrees of dose reduction or image quality improvement at the same dosecan be achieved. Therefore, for a given subject size and clinicalapplication, an optimal tube potential exists that yields the best imagequality or the lowest radiation dose.

Existing clinical studies have used empirically-determined tubepotentials for a certain patient group, with various levels of dosereduction or image quality improvement being observed. However, an exactknowledge of the dose-efficiency of different tube potentials to obtaina target image quality, and the dependence of the optimal tube potentialon patient size and diagnostic task, remain to be quantitativelydetermined. In a more recent study, some have used a dose-normalizedcontrast to noise ratio (CNR) as the criterion to determine the optimaltube potential and to quantify its dependence on phantom sizes andcontrast materials. Their results demonstrated that the selection oftube potential should to be adapted to the patient size and to thediagnostic task to a much higher degree than is common practice today inorder to further reduce the radiation dose. However, clinical practicesto prospectively select the optimal tube potential and determine thecorresponding radiation dose level that takes into account both thepatient size and the target image quality required by a particulardiagnostic task are lacking.

SUMMARY OF THE INVENTION

The present invention overcomes the aforementioned drawbacks byproviding a system and method that allows automatic adaptation of tubepotential as a function of patient size and diagnostic task and uses anovel image quality index, “noise-constrained iodine contrast to noiseratio,” to quantify the different levels of image quality suitable forvarious clinical applications.

In particular, the present invention provides a method for performingcomputer tomography (CT) imaging with automatic tube potentialselection. This method includes positioning a subject to be imaged in aCT system, selecting a scanning technique at a reference tube potentialcorresponding to a desired image quality level, and selecting a noiseconstraint parameter according to the diagnostic task. The methodfurther includes determining the dose efficiency of each tube potentialusing characteristics of the subject and noise constraint parameter, andperforming a CT scan using the most dose-efficient tube potential,provided that the system power limit is not exceeded and the scanningtime is practical.

Various other features of the present invention will be made apparentfrom the following detailed description and the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A and 1B depict a CT system configured to employ an automatictube potential selection method in accordance with the presentinvention; and

FIG. 2 is a flowchart setting for the steps for an automatic tubepotential selection method in accordance with the present invention.

DESCRIPTION OF THE PREFERRED EMBODIMENT

With initial reference to FIGS. 1A and 1B, the present invention may beimplemented using a computed tomography (CT) imaging system 10 thatincludes a gantry 12 and is representative of a “third generation” CTscanner, though the present invention can be implemented on anygeneration of CT scanner. Gantry 12 has an x-ray source 13 that projectsa fan beam or cone beam of x-rays 14 toward a detector array 16 on theopposite side of the gantry. The detector array 16 is formed by a numberof detector elements 18 which together sense the projected x-rays thatpass through a medical subject 15. Each detector element 18 produces anelectrical signal that represents the intensity of an impinging x-raybeam and hence the attenuation of the beam as it passes through thesubject. During a scan to acquire x-ray projection data, the gantry 12and the components mounted thereon rotate about a center of rotation 19located within the subject 15.

The rotation of the gantry and the operation of the x-ray source 13 aregoverned by a control mechanism 20 of the CT system. The controlmechanism 20 includes an x-ray controller 22 that provides power andtiming signals to the x-ray source 13 and a gantry motor controller 23that controls the rotational speed and position of the gantry 12. A dataacquisition system (DAS) 24 in the control mechanism 20 samples analogdata from detector elements 18 and converts the data to digital signalsfor subsequent processing. An image reconstructor 25, receives sampledand digitized x-ray data from the DAS 24 and performs high speed imagereconstruction. The reconstructed image is applied as an input to acomputer 26 which stores the image in a mass storage device 29.

The computer 26 also receives commands and scanning parameters from anoperator via console 30 that has a keyboard. An associated display 32allows the operator to observe the reconstructed image and other datafrom the computer 26. The operator supplied commands and parameters areused by the computer 26 to provide control signals and information tothe DAS 24, the x-ray controller 22 and the gantry motor controller 23.In addition, computer 26 operates a table motor controller 34 whichcontrols a motorized table 36 to position the subject 15 in the gantry12.

The present invention provides a method for selecting tube potential ina CT system to reduce radiation dose. Specifically, referring now toFIG. 2, an automatic tube potential selection method using the system,for example, of FIGS. 1A and 1B is provided that uses a dose efficiencyevaluation process to achieve dose reduction. The method begins atprocess block 202 with the acquisition of a scanned projectionradiograph, which provides a basis for both automatic tube currentmodulation and automatic tube potential selection.

In conventional automatic tube current modulation, the tube current ismodulated according to the attenuation level of the subject at eachprojection angle. The purpose of the modulation is to use appropriatetube current to generate a desired noise level in the CT images.Therefore, the noise level is an image quality index that is used inautomatic tube current modulation.

X-ray attenuation can be defined as A(d)=exp(μd) and a generalsimplified tube current modulation model that relates the tube currentto attenuation path length d can be given by:

$\begin{matrix}{{{I(d)} = {{I\left( d_{0} \right)} \cdot \left( \frac{A(d)}{A\left( d_{0} \right)} \right)^{x}}};} & {{Eqn}.\mspace{14mu} 1}\end{matrix}$

where I(d) and I(d₀) respectively denote the tube current forattenuation path length d and d₀, which is the reference attenuationpath length for which the reference tube current I(d₀). In addition, μdenotes the average linear attenuation coefficient along the path lengthand x is the tube current modulation strength. When x=1, a constantnoise σ(d₀) is maintained in each projection independent of path length.When x<1, the noise level is lower than σ(d₀) when the attenuation pathlength is less than the reference length d₀ and higher than a (d₀) whenthe attenuation path length is longer than the reference length d₀.

Noise in each set of projection data (indexed by i) after logarithmicoperation can be related to the incident number of photons andattenuation by:

$\begin{matrix}{{\sigma_{i}^{2} = \frac{A_{i}}{N_{0\; i}}};} & {{Eqn}\;.\mspace{11mu} 2.}\end{matrix}$

wherein N_(0i) is assumed to be proportional to the tube current.

Therefore, with the tube current modulation defined in Eqn. 1, the noisein projection data with an attenuation path length of d is given by:

$\begin{matrix}{{\sigma (d)} = {{\sigma \left( d_{0} \right)} \cdot {\left( \frac{A(d)}{A\left( d_{0} \right)} \right)^{{({1 - x})}/2}.}}} & {{Eqn}\;.\mspace{11mu} 3}\end{matrix}$

The relationship between image noise and the noise in the projectiondata is generally a complicated function that is dependent on thereconstruction algorithm and smoothing filter. This can be simplified byconsidering only the pixel noise in the center of the image, which canbe related to the noise in each projection as:

$\begin{matrix}{{{\sigma^{2}\left( \Omega_{d} \right)} = {\sum\limits_{i = 1}^{N}\; {\sigma^{2}\left( d_{i} \right)}}};} & {{Eqn}\;.\mspace{11mu} 4}\end{matrix}$

where Ω_(d)=(d_(i)|i=1, . . . , N) denotes a given subject size andd_(i) is the x-ray path length passing through the isocenter at the i-thprojection view. This provides a simplified representation of thesubject size. Using a simplified model, some concluded that the minimumimage noise is obtained when the modulation strength x is 0.5 for agiven total radiation dose, that is, a fixed total number of incidentphotons.

Contrast-to-noise ratio (CNR) may also be used as an image qualityindex. If the attenuation of the contrast material along each projectionremains the same at exp(μ_(c)d_(c)), then the contrast in sinogram spaceis C=(μ_(c)−μ_(w))d_(c), which is independent of the background pathlength d. Therefore, the CNR in the projection is given by:

$\begin{matrix}{{C\; N\; {R(d)}} = {\frac{\left( {\mu_{c} - \mu_{w}} \right)d_{c}}{\sigma (d)}.}} & {{Eqn}\;.\mspace{11mu} 5}\end{matrix}$

When the tube potential is fixed, the contrast is fixed. Therefore,using CNR as an image quality index to modulate the tube current isequivalent to using noise as the image quality index. When tubepotential can also be modulated for each projection view, the contrastis a function of tube potential, which is given byC(kV)=[μ_(c)(kV)−μ_(w)(kV)]d_(c). The CNR in a projection can thus beexpressed as:

$\begin{matrix}{{C\; N\; {R\left( {d,{kV}} \right)}} = {\frac{\left\lbrack {{\mu_{c}({kV})} - {\mu_{w}({kV})}} \right\rbrack d_{c}}{\sigma \left( {d,{kV}} \right)}.}} & {{Eqn}\;.\mspace{11mu} 6}\end{matrix}$

The case in which the CNR at different tube potentials is the same as ina reference tube potential for each path length d may be expressed asfollows:

CNR(d,kV)=CNR(d,kV _(ref))  Eqn. 7.

For each projection view, if the chosen tube potential is the mostdose-efficient one for achieving the desired CNR defined in thereference tube potential, then the total radiation dose is the least.This can be likened to automatic tube current modulation, where the tubepotential is dynamically modulated during the scan to maintain the samelevel of CNR for each projection view. However, this strategy does notwork well for a number of reasons when applied in a clinical setting.When tube potential is changing during a single scan, the attenuationcoefficients of each voxel element also changes as a function ofprojection angle. In this case, the map of attenuation coefficients,and, thus, the image trying to be reconstructed, becomes a function ofprojection angle and image reconstruction is made prohibitivelydifficult due to data inconsistencies. Unless there is a perfectbeam-hardening correction for each kV, tube potential modulation cannotbe implemented in the same way as tube current modulation. Thus, thestrategy of automatic tube potential modulation presented here willassume tube potential does not change in a single scan. However, thegeneral idea of determining the most dose efficient tube potentialaccording to the attenuation level and noise constraint parameter, whichis described below, will also work if the tube potential is modulated ateach projection angle.

A simplified representation of the subject size can be described byΩ_(d)={d_(i)|i=1, . . . , N}, where d_(i) is the x-ray path lengthpassing through the isocenter at the i-th projection view. The contrastof a particular material, for example, iodine, calcium, or soft tissue,relative to the background material is denoted as C(Ω_(d)), whichrepresents a function of contrast changing with the subject size. Thus,the CNR is given by:

$\begin{matrix}{{C\; N\; {R\left( \Omega_{d} \right)}} = {\frac{C\left( \Omega_{d} \right)}{\sigma \left( \Omega_{d} \right)}.}} & {{Eqn}\;.\mspace{11mu} 8}\end{matrix}$

To find the most dose efficient tube potential that gives the desiredCNR for a given subject size Ω_(d), the desired CNR as defined by thereference tube potential and radiation dose for subject sizeCNR(Ω_(d),kV_(ref),D_(ref)) may be considered. Thus, the tube potentialthat uses the lowest radiation dose to achieve the same value of CNR canbe expressed as:

$\begin{matrix}{\underset{kV}{argmin}{{D\left( {\Omega_{d},{kV},{C\; N\; R}} \right)}.}} & {{Eqn}\;.\mspace{11mu} 9}\end{matrix}$

This objective may also be expressed in another way by considering whichtube potential yields the minimum D under the following condition:

CNR(Ω_(d) ,kV,D)=CNR(Ω_(d) ,kV _(ref) ,D _(ref))  Eqn. 10.

This equation may be rewritten to give:

$\begin{matrix}{{\sigma \left( {\Omega_{d},{kV},D} \right)} = {\frac{C\left( {\Omega_{d},{kV}} \right)}{C\left( {\Omega_{d},{kV}_{ref}} \right)}{{\sigma \left( {\Omega_{d},{kV}_{ref},D_{ref}} \right)}.}}} & {{Eqn}\;.\mspace{11mu} 11}\end{matrix}$

Therefore, by assuming σ²(Ω_(d),kV,D)=k(Ω_(d),kV)/D, where k(Ω_(d),kV)is a coefficient that can determined experimentally, the following canbe determined:

$\begin{matrix}{D = {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2}\frac{k\left( {\Omega_{d},{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}{D_{ref}.}}} & {{Eqn}\;.\mspace{11mu} 12}\end{matrix}$

A relative dose factor (RDF) that represents the relative dose at a tubepotential can be defined. The RDF can also be related to dose efficiencyε for a specific subject size and a desired image quality index can bedefined therewith. For example, as shown above, when the CNR is used asan image quality index, RDF is defined to be:

$\begin{matrix}{{{R\; D\; {F\left( {\Omega_{d},{kV}} \right)}} = {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2}\frac{k\left( {\Omega_{d},{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}}};} & {{Eqn}\;.\mspace{11mu} 13}\end{matrix}$

Therefore, the dose for another tube potential to maintain the same CNRas in the reference tube potential is given by the expression:

D=RDF(Ω_(d) ,kV)D _(ref)  Eqn. 14.

With this in mind, it is possible to quantify the relative doseefficiency of a tube potential for an image quality index such as iodineCNR. If, for example, a reference tube potential is 120 kV, then for asubject of size Ω_(d), the RDF is 1.2 for 140 kV. This means that 140 kVinvolves a 20% higher radiation dose to get the same image quality as120 kV. Likewise, if RDF is 0.5 for 80 kV, then 80 kV involves 50% lessradiation dose to get the same image quality as 120 kV.

However, iodine CNR alone is not always a good image quality index, asthere are some clinical applications where it undesirable to maintain aconstant iodine CNR, for example, situations where noise is excessivelyhigh at a lower tube potential, even though contrast enhancement at thelower tube potential is significantly better than at a higher tubepotential. Therefore, a general image quality index may be defined asfollows so that an additional variable is provided to adjust image noiseaccording to clinical need:

CNR(Ω_(d) ,kV,D)≧CNR(Ω_(d) ,kV _(ref) ,D _(ref))

σ(Ω_(d) ,kV,D)≦ασ(Ω_(d) ,kV _(ref) ,D _(ref))  Eqn. 15;

where α can be adjusted to allow varying conditions of noise at thereference tube potential. This gives:

$\begin{matrix}{{D \geq {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2}\frac{k\left( {\Omega_{d}{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}D_{ref}}}{D \geq {\frac{k\left( {\Omega_{d},{kV}} \right)}{\alpha^{2}{k\left( {\Omega_{d},{kV}_{ref}} \right)}}D_{ref}}}{D = {\max {\left\{ {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2},\frac{1}{\alpha^{2}}} \right\} \cdot \frac{k\left( {\Omega_{d},{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}}{D_{ref}.}}}} & {{Eqn}\;.\mspace{11mu} 16}\end{matrix}$

Thus, the RDF at each tube potential for achieving a target imagequality can be defined as follows:

$\begin{matrix}{{R\; D\; {F\left( {\Omega_{d},{kV}} \right)}} = {\max {\left\{ {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2},\frac{1}{\alpha^{2}}} \right\} \cdot {\frac{k\left( {\Omega_{d},{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}.}}}} & {{Eqn}\;.\mspace{11mu} 17}\end{matrix}$

Referring again to FIG. 2 and particularly, at process block 204, areference tube potential and a CT scanning technique is selected. Forexample, on a Siemens scanner this may include the selection of aquality reference mAs (QRM) if automatic tube current modulationCAREDose4D is active. On a GE scanner, this may include the selection ofa noise index if AutomA or SmartmA is active. According to theattenuation information provided by the scanned projection radiograph,the scanner can automatically determine the tube current for eachprojection angle and scan location and obtain the total radiation doseas expressed in CTDI_(vol). If automatic tube current modulation isinactive, then a constant tube current may be used for the scan.

At process block 206, a noise constraint parameter α is selected. It iscontemplated that the noise constraint parameter is based on the task ofthe CT scan being performed. For example, a CT angiogram may involve aniodine CNR with a wide noise constraint, such as α=1.5-2.0, whichimplies that the iodine CNR is equivalent or improved and that the noiseis allowed to be 50-100 percent higher than the reference level. For acontrast-enhanced chest, abdomen, and pelvis examination, an iodine CNRwith a moderate noise constraint, such α=1.1-1.25, may be used. For anon-contrast chest, abdomen, and pelvis examination an iodine CNR with atight noise constraint, such as α=1.0, may be used and, for CTenterography an iodine CNR with a noise constraint α=1.2-1.5 may beused.

At process block 208, RDF is generated based on scan parameters such asthe attenuation level and the selected noise constraint parameter. Forexample, using Eqn. 17 the RDF is calculated for a range of tubepotential's based on subject size and the selected noise constraintparameter and reference tube potential. The contrast and noiseinformation, expressed by C(Ω_(d),kV) and k(Ω_(d),kV), respectively, canbe determined using mathematical models or lookup tables generated fromphantom measurement. Generally, the most dose-efficient tube potentialfor achieving a desired image quality is given by the tube potentialwith the smallest RDF. Thus, at process block 210, the tube potentialproviding the highest relative dose efficiency, that is, the tubepotential having the smallest RDF, is selected and the associatedradiation dose is determined by multiplying the original CTDI_(vol) atthe reference tube potential by RDF. An mA setting (mA, mAs, effectivemAs, noise index, or quality reference mAs) that yields the CTDI_(vol)at each tube potential is subsequently calculated at process block 212.

At decision block 214, the CT system parameters and scanning time arechecked to determine if the dose level is practically achievable usingthe selected tube potential. If so, a scan using this tube potential isperformed at process block 216 and the scan quality reference tube mAs(QRM) or noise index may be adjusted to provide a desired radiation doselevel if automatic tube current modulation is used. If, at decisionblock 214, the dose is not achievable, then the tube potential with thenext highest relative dose efficiency, that is, the next smallest RDFvalue, is selected at process block 218 and its corresponding tubecurrent is calculated at process block 212. If, at decision block 214,this tube potential value is deemed acceptable, then a scan is performedat process block 216 as described above. Otherwise, this cycle continuesuntil an appropriate tube potential is determined.

Thus, the present invention provides a method that directly quantifiesthe relative dose efficiency at each tube potential and provides astraightforward mechanism for selecting optimal tube potential based onthe calculated dose savings, or dose waste, for each tube potential.This allows straightforward and effective dose reduction based onsubject size and other scan parameters.

The present invention's generalized image quality index “iodine CNR withnoise constraint,” takes into account the various image qualityrequirements of different diagnostic tasks. RDF is generally dependenton the image quality index used to define the target image quality. Aswith conventional automatic tube current modulation where noise servesas the image quality index and the tube current is modulated to obtainthe desired noise level for each attenuation level, an appropriate imagequality index is also needed for automatic tube potential selection.Since iodine is the commonly used contrast media in CT, iodine CNR istypically used, either implicitly or explicitly, in existing studies atlower tube potential. In clinical practice, iodine CNR may generallyprovide a good indicator of image quality for CT angiograms in whichiodine-enhanced vessels are of primary interest. However, iodine CNRalone is not appropriate for every diagnostic task. For example, incontrast-enhanced chest/abdominal CT exams, a contrast agent may helpvisualize some important structures, but provides little contrastenhancement for other structures of interest. In this case, it isgenerally beneficial to constrain image noise. Using iodine CNR alone asan image quality index for selecting an optimal tube potential in suchexams may potentially yield images with excessive image noise andreduced diagnostic confidence. The “iodine CNR with noise constraint”generalized image quality index incorporates such considerations so thatnoise constraint parameters provides a flexible means of characterizingand accommodating the various image quality requirements of differentdiagnostic tasks.

In general, the noise constraint parameter a can be used to control thenoise level allowed for a new tube potential relative to the referencetube potential. With a tight noise constraint close to 1, the noiseshould be approximately matched for different tube potential selectionsand there is limited potential for dose saving at lower tube potentialvalues. With a loose noise constraint, for example, α=2, a significantlylarger noise level is permitted while the same CNR is maintained due toimproved iodine enhancement. Therefore, the radiation dose can bereduced considerably for smaller subject sizes. It is contemplated thatnoise constraint selection is dependent on clinical application. Foriodine-enhanced angiographic study, the vessel visibility is of primaryinterest. Since the vessel is enhanced better at lower tube potential byiodine contrast, a relatively higher noise level can be tolerated andrelatively high noise constraint can be used. Alternately, while theconspicuity of hypo- or hyper-vascular structures is better at lowervalues of tube potential in iodine-enhanced chest and abdomen CT exams,many other structures with reduced iodine uptake are still of diagnosticinterest. Therefore, excessive noise should not be allowed and arelatively tight noise constraint should be applied. The appropriatevalue of noise constraint parameters for other diagnostic tasks shouldbe carefully evaluated in clinical practice.

When clinically implementing automatic tube potential selection, inaddition to selecting an appropriate noise constraint parameter, it isbeneficial to select a reference quality mAs level or noise index at thereference tube potential that is appropriate for the diagnostic task athand. Such steps help avoid excess radiation dose and image noise. Theseselections can be set up beforehand for each scanning protocol. Based onparameters provided by a user, the automatic tube potential selectionmethod provided by the present method allows the most dose-efficienttube potential to be selected for a given scan.

The present invention provides a system and method to directly quantifythe relative dose that is required at each tube potential to achieve atarget image quality. Just as with AEC systems, where noise level canserve as the image quality index and the tube current is modulated toobtain the desired noise level, an appropriate image quality index alsoneeds to be defined for automatic tube potential selection. Iodine CNRcan be used as the image quality index, since iodine is so commonly usedas a contrast agent in CT exams. However, it is generally not anappropriate index for all diagnostic tasks. For example, in routinecontrast-enhanced chest/abdominal CT exams, although the increasedcontrast enhancement at lower-tube potential is helpful for visualizingsome important structures, there is little enhancement in otherstructures that are also of interest in the diagnosis. Therefore, imagenoise cannot be too high in these type of clinical application. Usingiodine CNR alone as an image quality index for selecting optimal tubepotential will yield images with too high image noise and jeopardize thediagnostic confidence. Thus, a generalized image quality index, iodineCNR with a noise constraint, is defined in the accordance with thepresent invention to accommodate different diagnostic tasks that havedifferent image quality requirements. Thus, iodine CNR is a special caseof this general quality index. The noise constraint parameter can becontinuously adapted for different clinical applications.

The present invention has been described in terms of the preferredembodiment, and it should be appreciated that many equivalents,alternatives, variations, and modifications, aside from those expresslystated, are possible and within the scope of the invention. Therefore,the invention should not be limited to a particular describedembodiment.

1. A method for performing computer tomography (CT) imaging withautomatic tube potential selection, the method comprising the steps of:disposing a subject to be imaged in a CT system; selecting a scanningtechnique to be performed using at least one of a reference tubepotential and a tube-current-time-product corresponding to a selecteddose level and a desired image quality level; selecting anoise-constraint parameter to define a target image quality related to adesired contrast to noise ratio (CNR) for the scanning technique;determining a tube potential for the CT system and an associated doseefficiency using a characteristic of the subject, the diagnostic task,and the noise-constraint parameter; and performing a CT scan using thescanning technique and tube potential with a dose-efficiency selected toachieve the target image quality.
 2. The method as recited in claim 1further comprising performing a scout scan using the CT system, whereinat least one of the scanning technique at the reference tube potentialand a tube current for the CT scan is selected based on thecharacteristic of the subject as determined by the scout scan.
 3. Themethod as recited in claim 1 wherein the characteristic is at least oneof a size property and attenuation property.
 4. The method as recited inclaim 1 wherein determining the tube potential for the CT system withthe desirable achievable dose efficiency includes: generating a relativedose efficiency for a range of tube potentials based on thecharacteristic of the subject and the selected reference tube potentialand noise-constrained, contrast-to-noise parameter; and selecting a tubepotential with a most desirable relative dose efficiency.
 5. The methodas recited in claim 4 wherein generating the tube potential doseefficiency employs the relationship:${R\; D\; {F\left( {\Omega_{d},{kV}} \right)}} = {\max {\left\{ {\left\lbrack \frac{C\left( {\Omega_{d},{kV}_{ref}} \right)}{C\left( {\Omega_{d},{kV}} \right)} \right\rbrack^{2},\frac{1}{\alpha^{2}}} \right\} \cdot \frac{k\left( {\Omega_{d},{kV}} \right)}{k\left( {\Omega_{d},{kV}_{ref}} \right)}}}$which can also be expressed as:${{R\; D\; {F\left( {\Omega,{kV}} \right)}} = {\left\{ {\min \left( {\frac{C\left( {\Omega,{kV}} \right)}{C\left( {\Omega,{kV}_{ref}} \right)},\alpha} \right)} \right\}^{2} \cdot \frac{k\left( {\Omega,{kV}} \right)}{k\left( {\Omega,{kV}_{ref}} \right)}}};$in which RDF(Ω_(d),kV) denotes a relative dose at a tube potential kV(relative to the dose at a reference tube potential kV_(ref)),C(Ω_(d),kV) and C(Ω_(d),kV_(ref)) denote contrast levels, k(Ω_(d),kV)and k(Ω_(d),kV_(ref)) denote coefficients relating image noise levels todose levels, α denotes the noise-constraint parameter, Ω_(d) denotes thecharacteristic of the subject, and kV_(ref) denotes the reference tubepotential.
 6. The method as recited in claim 5 wherein the tubepotential with the highest relative dose efficiency is a tube potentialwith a smallest value of RDF(Ω_(d),kV).
 7. The method as recited inclaim 4 wherein determining the tube potential for the CT system withthe desirable achievable dose efficiency further includes: estimating anexpected radiation dose for the selected tube potential; and determiningif the expected radiation dose is achievable based on parameters of theCT system and scanning time.
 8. The method as recited in claim 7 whereindetermining if the expected radiation dose is achievable includes:calculating a tube current necessary to provide the expected radiationdose; and determining if the calculated tube current is achievable bythe CT system within a selected scanning time.
 9. The method as recitedin claim 7 further comprising selecting a tube potential with anincreased relative dose efficiency if the tube potential with thedesirable relative dose efficiency is unachievable.
 10. The method asrecited in claim 1 wherein the target image quality requires that theCNR and noise in associated with target image quality using the tubepotential and satisfy the following two conditions:CNR≧CNR _(ref) & σ≦ασ_(ref), where CNR_(ref) and σ_(ref) denote CNR andthe noise level obtained in a reference tube potential and dose level,CNR and σ denote CNR and noise level obtained with target image qualityusing the tube potential, and α denotes a noise constraint definedaccording to the diagnostic task.
 11. The method as recited in claim 5wherein the two conditions to define the target image quality isequivalent to:${\sigma \leq {{\min \left( {\frac{C\left( {\Omega,{kV}} \right)}{C\left( {\Omega,{kV}_{ref}} \right)},\alpha} \right)} \cdot \sigma_{ref}}},$where the noise constraint parameter α is equivalent to a constraint oncontrast ratio between two tube potentials.
 12. The method as recited inclaim 1 wherein determining the tube potential for the CT system withthe desirable achievable dose efficiency includes determining a relativedose at each tube potential for achieving a target image quality definedby a reference scan technique and the CNR.
 13. The method of claim 14wherein determining the tube potential for the CT system includesattempting to achieve the target image quality using radiation doseefficiency over a range of tube potentials and based on at least one ofsubject size, image quality, and noise constraints.
 14. The method asrecited in claim 1 wherein automatic tube current modulation is employedfor the CT scan and at least one of a quality reference tubecurrent-time-product or a noise index is adjusted to provide a desiredradiation dose level.
 15. A computer tomography (CT) imaging systemconfigured to: determine at least one of a reference tube potential anda tube-current-time-product corresponding to a selected dose level and adesired image quality level for a prescribed medical imaging procedureof a subject to be performed using the CT imaging system; determine anoise-constraint parameter associated with the prescribed medicalimaging procedure of the subject to be performed; determine a tubepotential for the prescribed medical imaging procedure to be performedand an associated desired dose efficiency using at least one of acharacteristic of the subject, the reference tube potential, thetube-current-time-product, and the noise-constrained CNR; and performthe prescribed medical imaging procedure using the CT imaging system atthe tube potential and based on the desired achievable dose efficiency.